Wayne State University Wayne State University Theses
1-1-2015
Methodology For Performing Whole Body Pmhs
Underbody Blast Impact Testing, And The
Corresponding Response Of The Hybrid Iii
Dummy And The Finite Element Dummy Model
Under Similar Loading Condition
Karthik Somasundaram
Wayne State University,
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Recommended Citation
Somasundaram, Karthik, "Methodology For Performing Whole Body Pmhs Underbody Blast Impact Testing, And The Corresponding Response Of The Hybrid Iii Dummy And The Finite Element Dummy Model Under Similar Loading Condition" (2015). Wayne State University Theses. 535.
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Methodology for Performing Whole Body PMHS Underbody Blast Impact Testing, and the Corresponding Response of the Hybrid III Dummy and the Finite Element
Dummy Model under similar Loading Condition
By
Karthik Somasudaram THESIS
Submitted to the Graduate School Of Wayne State University,
Detroit, Michigan
In partial fulfillment of the requirements For the degree of
MASTER OF SCIENCE 2016
MAJOR: BIOMEDICAL ENGINEERING (Impact Biomechanics)
Approved By: __________________ ______________________________
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©COPYRIGHT BY KARTHIK SOMASUNDARAM
2016
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ACKNOWLEDGEMENTS
This study was funded by contract #N00024-13-D-6400, U.S. Army Research, Development and Engineering Command. Additional support was awarded by DOD, Army Research Office, National Defense Science and Engineering Graduate (NDSEG) Fellowship, 32 CFE 168a and Supporting Universities.
I would like to take this opportunity to thank all those people who have made an immense contribution towards the completion of my master’s thesis at Wayne State University.
I feel fortunate and thankful to have Dr.Cavanaugh as my advisor, for his valuable support, guidance and time throughout my thesis and research work. Moreover, his endless patience and encouragement have propelled towards accomplishing my thesis. I am grateful to Dr. Cavanaugh to show confidence in me and giving me the opportunity to conduct a collaborative study with other laboratories.
I would also like to thank my committee members, Dr. Paul Begeman and Dr. Albert King, for providing their valuable motivation and direction towards completion of my thesis.
Further, I would like to express my sincere gratitude towards Mr. Don Sherman and Mr. Jason Greb for their guidance and the Sport & Ballistics group of the Biomedical Engineering Department at Wayne State University for assisting me in research work. I extend my gratitude to the Advance Human Modeling lab, especially to Mr. Anil Kalra, who helped me to understand the FEM concept and assisted me in running simulations.
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Further Dr. Feng Zhu’s expertise in FEA was essential to the completion of the simulation studies. I appreciate Dr. Cavanaugh’s spine laboratory group for their timely assistance in this research and thesis completion work. I am very grateful to Dr. Srinivasu Kallakuri for motivating and encouraging me throughout my graduate study at Wayne State University.
I appreciate the effort put by Darlene Pennington-Johnson and Loretta Ann Sabo, Office of Research Compliance, for providing the necessary permission and license from the State Department to work with the sensor in this project. Special thanks go to radiologist Dr.Kochkodan for examining CT scan data, Dr.Schmidt for conducting autopsy evaluations, and the orthopedic residenct for assisting with the spine instrumentation.
Last I express my thanks to family members, my Mom, Dad, grandfather and all of my friends, for their unconditional love and support during the duration of my master’s thesis.
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TABLE OF CONTENTS
ACKNOWLEDGEMENTS ... iii
TABLE OF CONTENTS ... v
LIST OF FIGURES ... viii
LIST OF TABLES ...xiii
CHAPTER 1- INTRODUCTION ... 1
1.1 Aim of the study ... 3
1.2 Epidemiology of blast-related skeletal injuries ... 4
1.3 Positioning of the occupant ... 10
CHAPTER 2- THEATER INJURY CASE STUDY AND BLAST BIOMECHANICS ... 14
2.1 Case study from literature ... 14
2.2 Effect of blast event on a vehicle and its occupant ... 16
CHAPTER 3- ANATOMY AND BONE MINERAL DENSITY ... 19
3.1 Spine ... 19
3.2 Pelvis ... 21
3.2.1 Sacrum ... 22
3.3 Ligaments of spine and pelvis ... 24
3.4 Femur ... 25
3.5 Bone mineral density(BMD) ... 26
CHAPTER 4- CADAVER TESTING AND DATA PROCESSING METHODS ... 31
4.1 Horizontal sled system ... 32
4.2 Data acquisition and camera setup ... 34
4.3 Loading Condition ... 35
4.4 PMHS preparation ... 37
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4.4.2 Instrumentation ... 39
4.5 Specimen positioning ... 45
4.5.1 Positioning Protocol ... 46
4.6 Pre-test sled preparation ... 50
4.7 Post-impact procedure... 51 4.8 Data processing ... 52 4.8.1 Converted Staging ... 53 4.8.2 Processed staging ... 53 4.8.3 Calculated staging ... 55 4.8.4 Scaled Staging ... 62
CHAPTER 5- CADAVER IMPACT RESPONSE ... 67
5.1 Impact Response ... 68
5.1.1 Spinal Compliance ... 96
5.2 Video Kinematic Analysis ... 97
CHAPTER 6- AN EXPERIMENTAL AND NUMERICAL STUDY OF THE HYBRID III DUMMY’S RESPONSE TO SIMULATED UNDERBODY BLAST IMPACTS ... 106
6.1 Introduction ... 106
6.2 Methodology ... 107
6.2.1 Data Acquisition System and Dummy Instrumentation ... 108
6.2.2 Simulation Setup ... 109 6.3 Results ... 113 6.4 Discussion ... 124 CHAPTER 7- CONCLUSION ... 130 8 APPENDIX A ... 132 8.1 ANTHROPOMETRIC MEASUREMENT [cm] ... 132
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8.1 SEATED MEASUREMENT [cm] ... 133
8.2 ROMER ARM DATA ... 133
9 APPENDIX B ... 134
9.1 PMHS Instrumentation Channel Assignment Matrix ... 134
9.2 DATA PROCESSING MATRIX ... 137
10 APPENDIX C ... 138
10.1 WSU-003 SCALE FACTOR ... 138
10.2 WSU-004 SCALE FACTOR ... 138
11 APPENDIX D ... 139
11.1 ORIENTATION OF MOUNT RELATIVE TO THE BONE ... 139
12 APPENDIX E ... 140
13 ABSTRACT ... 14141 14 REFERENCES ... 1413 15 AUTOBIOGRAPHY……….…148
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LIST OF FIGURES
Figure 1-1: Pie chart highlights the percentage of USA military loss due to landmine explosions in past
warfare (Bird, 2001). ... 4
Figure 1-2: Contribution of different causative agents towards extremities injuries caused during OIF & OEF (Owens et al., 2007). ... 5
Figure 1-3: Axial compressive force measured in the tibia, lumbar spine and upper neck for a vertical load caused by underbelly blast impact (NATO, 2007)... 6
Figure 1-4: Pictorial representation of the load transmission paths subjected to vertical load under a UBB event (Ramasamy et al., 2009). ... 7
Figure 1-5: Pie chart of the percentage of pelvic injury caused by different weapons (Bailey et al., 2011). 9 Figure 1-6: Bar chart of the percentage of spine trauma in warfare injuries. *BCT- Brigade combat team deployed in OIF (Schoenfeld et al., 2012a). ... 10
Figure 1-7: Shift in the body’s center of gravity based on occupants sitting posture (Schoberth and Hegemann, 1962). ... 12
Figure 1-8: Effect of the thigh-torso angle on pelvic tilt (KEEGAN and Omaha, 1953). ... 13
Figure 3-1: A pictorial comparison between the lumbar and thoracic vertebral bodies (Gray et al., 1973). ... 21
Figure 3-2: Anatomy of the Spine (Gray et al., 1973). ... 21
Figure 3-3: Anatomy of the Pelvis (Gray et al., 1973). ... 22
Figure 3-4: Sacrum frontal view (Gray et al., 1973). ... 23
Figure 3-5: Anterior and posterior view of the femur (Gray et al., 1973). ... 26
Figure 3-6: Active duty military and civilians by race and ethnicity, 2002 (Segal and Segal, 2004). ... 27
Figure 4-1: A lateral view of the horizontal sled system with the occupant positioned on the seat with a 5-point belt. ... 32
Figure 4-2: Two large capacity snubbers mounted to the barrier to slow down the WHAM III. Four pre-crushed aluminum honeycomb blocks attached to the rigid barrier to produce short duration seat acceleration. ... 34 Figure 4-3: Structure acceleration and corresponding velocity curves are shown in sequence (A: WSU-003 Acceleration (7270 accelerometer), B: WSU-WSU-003 Velocity, C: WSU-004 Acceleration (7270
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accelerometer), D: Test WSU-004 Velocity, E: WSU-003 Acceleration (LoFFI accelerometer) and F:
WSU-004 Acceleration (LoFFI accelerometer)). ... 37
Figure 4-4: The head and sternum steel plate [A], and the tibia and femur mounts [B], respectively. ... 41
Figure 4-5: The spine mounts fabricated with different depths [A] and Individual spine mount assembly [B, C, and D]. ... 43
Figure 4-6: Sacrum mounts fabricated with different depth [A]. The Individual sacrum mount assembly [B]. ... 43
Figure 4-7: A 6DX sensor mounted at the right distal femur [A] and T5 vertebra [B], respectively. ... 44
Figure 4-8: Strain gauge installation sequence. ... 45
Figure 4-9: Pelvis angle measured by the Romer arm for Tests 3 and 4, respectively. ... 49
Figure 4-10: A- Final specimen position for test 4 loading conditions prior to impact. B- The contact between the foot/boot and boot/floor plate, respectively. C- The buttock contact with the seat bottom. Also, the tail end of the sacrum is observed in the same image. D- The knee angle between the femur and tibia. Along with the angle verification, the orientation of the 6DX block on anatomical landmarks could also be examined using lateral X-ray images. ... 50
Figure 4-11: Summarizes the procedure followed to calcuate the hip-joint center... 62
Figure 5-1: Illustration of the CT and autopsy demonstrating the impaction fracture of the L1 spine sustained by the OSU 6908 specimen. The test subject was impacted at a seat velocity of 4m/s with 10ms time to peak. ... 67
Figure 5-2: Illustration of the CT demonstrating the compression fracture of the L3 spine sustained by the LMD 14-00355 specimen. This test subject was also impacted at a seat velocity of 4m/s with 10ms time to peak. ... 68
Figure 5-3: Calcaneus Z Linear Acceleration ... 70
Figure 5-4: Tibia Z Linear Acceleration ... 71
Figure 5-5: Tibia XZ Resultant Acceleration ... 72
Figure 5-6: Tibia Y Rotation ... 73
Figure 5-7: Femur X Linear Acceleration ... 74
Figure 5-8: Femur XZ Resultant Acceleration ... 74
Figure 5-9: Femur Y Rotation ... 75
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Figure 5-11: Sacrum XZ Resultant Acceleration ... 77
Figure 5-12: Sacrum X Rotation ... 79
Figure 5-13: Sacrum Y Rotation ... 79
Figure 5-14: Sacrum Z Rotation ... 80
Figure 5-15: T12 Spine Z Linear Acceleration ... 81
Figure 5-16: T12 Spine XZ Resultant Acceleration ... 82
Figure 5-17: T8 Spine Z Linear Acceleration ... 83
Figure 5-18: T8 Spine XZ Resultant Acceleration ... 83
Figure 5-19: T5 Spine Z Linear Acceleration ... 84
Figure 5-20: T5 Spine XZ Resultant Acceleration ... 85
Figure 5-21: T1 Spine Z Linear Acceleration ... 86
Figure 5-22: T1 Spine XZ Resultant Acceleration ... 86
Figure 5-23: Thoracic Spine Y Rotation ... 87
Figure 5-24: WSU-003 Spine Z Linear Acceleration. ... 88
Figure 5-25: WSU-004 Spine Z Linear Acceleration. ... 89
Figure 5-26: Sternum Z Linear Acceleration ... 90
Figure 5-27: Sternum XZ Resultant Acceleration ... 90
Figure 5-28: Sternum Y Rotation... 91
Figure 5-29: Head Z Linear Acceleration ... 92
Figure 5-30: Head XZ Resultant Acceleration... 93
Figure 5-31: Head X Rotation ... 94
Figure 5-32: Head Y rotation ... 94
Figure 5-33: Head Z Rotation ... 95
Figure 5-34: Relative Motion ... 96
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Figure 5-36: Kinematic comparison between WSU-003 and WSU-004 at different time points which
includes: Time zero, 20, 100, 300 ms, respectively. ... 99
Figure 5-37: Target marker displacement (cm) in Z from Time zero to 300ms for WSU-003 [top] and WSU-004 [bottom]. Targets were placed at the lateral malleolus, boot heel, tragion, infraorbital notch, lateral epicondyle and proximal humerus, respectively in this study. ... 100
Figure 5-38: WSU-003 Target Marker Displacement along X [top] and Z [bottom]. ... 101
Figure 5-39: WSU-004 Target Marker Displacement along X [top] and Z [bottom]. ... 102
Figure 5-40: The Shoulder Motion. ... 103
Figure 5-41: The Right Knee Motion. ... 104
Figure 5-42: The Right Foot Motion Relative to the Right Tibia. ... 105
Figure 6-1: A, B- A rigid foot floor plate and four hydraulic shock absorbers used for test condition 1 with the same loading condition at the seat and foot floor. ... 108
Figure 6-2: A- Comparison between the finite element and physical Hybrid III sled setups. B, C – The pelvis and foot of the FE model and physical ATD contact with the seat and floor plate, respectively. .. 111
Figure 6-3: Loading condition 1 floor and seat acceleration curve for the five consecutive tests, respectively. ... 114
Figure 6-4: Loading condition 2 floor and seat acceleration curve for the five consecutive tests, respectively. ... 114
Figure 6-5: The pelvis acceleration for loading condition 1 and 2, respectively. ... 115
Figure 6-6: The chest acceleration for loading condition 1 and 2, respectively. ... 115
Figure 6-7: The pelvis acceleration for loading condition 1 and 2, respectively. ... 116
Figure 6-8: The upper neck force for loading condition 1 and 2, respectively. ... 116
Figure 6-9: The tibia force for loading condition 1 and 2, respectively. ... 116
Figure 6-10: The lumbar spine load for loading condition 2. ... 117
Figure 6-11: Snapshot showing the punctured pelvis flesh impacted with loading condition 2. After the first experiment the rupture worsen with additional tests. ... 118
Figure 6-12: Comparison of the WSU Hybrid III dummy model-predicted and physical test measured impact response for impact condition 2. ... 121
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Figure 6-13: Comparison of the Hybrid III kinematics and validated dummy model for impact condition 2. From top to bottom the frame represents Tzero, 10 ms, 30 ms, and 60 ms, respectively. ... 123 Figure 6-14: An isometric view of the different components of the FE pelvis model with the pelvis flesh and foam removed from the right side. ... 124 Figure 6-15: Side view of finite element pelvis captured at t = 0 ms and t = 11 ms, respectively. ... 125 Figure 6-16: The location of pelvis flesh rupture observed in the post test Hybrid III dummy for loading condition 2 coincided with the location where the FE model predicted maximum principal stresses. Below image shows the corresponding pelvis foam stress map. ... 126 Figure 6-17: Comparisons of the Hybrid III Test 1 relative response with consecutive test data, separated by each loading condition. ... 127
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LIST OF TABLES
Table 1-1: Injury location for wounded personnel due to different causative agents during OIF-I (Zouris et
al., 2006). ... 5
Table 1-2: The orientation of the pelvis, corresponding CG and body weight transferred with respect to sitting posture (Schoberth and Hegemann, 1962). ... 12
Table 3-1: A brief summary of pelvis-floor ligament anatomical location and physiological function. (These ligaments have been hypothesized to play a critical role in pelvis fracture mechanism under, a UBB type environment) (Gray et al., 1973; Tile et al., 2003). ... 25
Table 3-2: Active –service duty officers by rank, service and race/ethnicity (Segal and Segal, 2004). .... 27
Table 3-3: World Health Organization’s definitions based on Bone Density Levels (T-score) (NIH, 2012). ... 28
Table 3-4: Osteoporosis attribution probability by fracture type, race/ethnicity and age (Melton et al., 1997). ... 29
Table 4-1: PMHS impact test measured input parameters. ... 36
Table 4-2: Specimen matrix ... 37
Table 4-3: Rationale for performing CT scans. ... 39
Table 4-4: Cadaver impact test instrumentation matrix. ... 39
Table 4-5: A summary of the anatomical landmark and steel mount installation technique for each body region, respectively. ... 41
Table 4-6: Transverse and sagittal angle the thoracic spine pedicle angle measurement. ... 42
Table 4-7: Anatomical locations for motion target markers (J Ruppa, 2013). ... 49
Table 4-8: Summarized anatomical loandmark required to define the local bone coordinate system. ... 57
Table 4-9: Threshold value used for CT mass measurement. ... 65
Table 5-1: A Summary of the Peak Acceleration and the Corresponding Peak Velocity for Each Test. ... 69
Table 5-2: Comparison of the WSU-003 and WSU-004 measured pelvis and corresponding peak response experienced by the sacrum and T12 vertebrae, respectively. ... 78
Table 6-1: Experimental impact test condition matrix. ... 108
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Table 6-3: Material Model Matrix. ... 113 Table 6-4: The CORA Analysis Parameter used in this study. ... 113 Table 6-5: Quantitative comparison of numerical model predicted response with the experimental peak pelvis acceleration and corresponding FE model cross correlation rating for WSU model. ... 121 Table 6-6: Quantitative comparison of numerical model predicted response with the experimental peak responses for different body regions and corresponding FE model cross correlation rating for WSU model. ... 122
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CHAPTER 1
1 INTRODUCTION
Since World War I, the use of tanks as well as armored and infantry vehicles have increased. They better protect the occupants from enemies’ fire and provide feasibility to maneuver through rough terrain. Possley et al. (2012) analyzed the probability of injuries sustained by mounted and dismounted soldiers during the Iraq war (2001-2009). The authors selected 1,890 spinal trauma causalities for their investigation. They reported that 26% of mounted soldiers sustained spinal fractures whereas; the remaining injuries were contributed by dismounted soldiers. This indicates that a maneuver with an infantry vehicle in a live theater is safer compared to foot movement. The efficiency of the tank was improved by adding armor and advanced weapons. As a tactical measure, improvised explosive devices (IED), anti-tank (AT) and anti-vehicular (AV) landmine weapons were developed. These mines and IEDs are usually buried in a vehicle’s pathway and are detonated by sensing the vibration or the mass of the vehicle (Schneck, 1998). The explosion of the improvised explosive (IED) device has been reported to damage the integrity of the vehicle. IED detonation has contributed to significant loss to the military, both in terms of cost and human resources (Bird, 2001; Owens et al., 2007; Wang et al., 2001; Zouris et al., 2006). Bird (2001) reported that the percentage loss increased from 22% in World War II to 60% in the Somalia war. IEDs/landmines, the future warfare weapon produces severe damage to both the vehicle and its occupants. Radonic et al. point out in their article that mines exploded ¼ of the German tanks during the Russian-German war (Radonic et al., 2004). The advancement of armor used in infantry vehicles has led to the development of more lethal IEDs. These modernized weapons designed to maximize the energy along the occupant -Z axis is directed from toe to head (2014 SAE standard). Therefore, the vehicle becomes immobilized leading the
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detonation energy to concentrate along the Z axis (Mckay, 2010; Schneck, 1998). IED blast energy disintegrates the structure of the vehicle and is capable of producing high vertical acceleration. In the Rhodesian war (1972-1980) about 1409 vehicles were detonated by landmines, resulting in 632 deaths and 4410 injuries (Bird, 2001). Alvarez reported that of the 608 live theater causalities, 456 were caused by wound in action, while the remaining was due to killed in action (Alvarez, 2011). The author adds to say that in both sets fractures caused the most casualties, accounting for 53% compared to the causalities due to internal organ and concussion injuries. The US military is more concerned about the safety of the occupant than for the structural integrity of the vehicle. Hence, they are working closely with vehicle designer to device some strategies to mitigate these injuries. IED detonations under the vehicle produce two vertical impulses, one at the feet and another at the buttocks of the occupant. Several lower extremity UBB impact studies have been performed in recent years (Bir et al., 2008; Mckay, 2010; Schueler et al., 1995; Van der Horst et al., 2005). These experiments and live fire studies have provided a better understanding about lower leg injury mechanisms. Further biomechanical study is required to understand the mechanism of injury and tolerance to skeletal structure exposed to UBB. Next, during a UBB event along with seat acceleration, the load from the tibia is predicted to transfer to the lower vertebral column. Hence, the spine and the pelvis are more vulnerable to this kind of impact. Automotive crashes have never yielded such a complex injury pattern; however, free fall trauma, parachute jumping and pilot seat ejection events give some insight into understanding IED associated vertical deceleration injury mechanism. The peak acceleration of a pilot seat ejection event’s falls in the range of 140 to 160 m/s2 (Miller and Morelli, 1993). A helicopter vertical load generates peak acceleration in the range of 320 to 400 m/s2 (Jackson et al., 2004) whereas the peak acceleration in a UBB blast is determined to be
3 980m/s2
(Wang et al., 2001). Unlike the events mentioned above, a UBB impact produces high kinetic energy over a couple of milliseconds predominately along the principal Z axis, making it unique in producing complex vertical deceleration injuries.
1.1 Aim of the study
The current study provides a detailed overview of the methodology for performing underbody blast impact testing in a laboratory environment. In addition, Hybrid III dummy response to vertical loading conditions caused by UBB impacts was investigated. Furthermore, Livermore Software Technology Corporation (LSTC) finite element model was updated and then validated against the Hybrid III test data. The following are the specific areas this study seeks to clarify.
A brief overview of orthopedic injuries due to an underbody blast (UBB) event
Literature review of the lower spine and pelvis injuries among soldiers in modern warfare subjected to IED detonation as well as a brief summary of UBB associated orthopedic injury conducted under laboratory set up
An overview of the spine and pelvis anatomy and bone mineral density measurement A detailed description of the cadaver testing methodology and data processing techniques
implemented for simulated underbody blast impact test in the WIAMan project at Wayne State University.
A report on the mechanical responses and the injuries produced from two postmoterm human surrogate (PMHS) tests for 4 m/s; 10 ms at the seat and 6 m/s; 5 ms at the floor.
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Develop and validate a finite element Hybrid III dummy in response to two simulated vertical loading conditions.
1.2 Epidemiology of blast-related skeletal injuries
The statistical analysis of soldier casualties reports that IED explosions, suicides, and roadside bombings have accounted for 60% loss of human life in Iraq and 50% in Afghanistan (Wilson, 2006). Landmine and IED explosions contribute to these losses and becoming a threat to military operations as well as human life. Different warfare statistics shown in Figure 1-1 highlight the percentage of USA military loss due to landmines in past combat operations (Bird, 2001). Owens et al. (2007) in their paper presented statistics about the contribution of modern warfare weapons to injuries during operations in Iraq (Operation Iraqi Freedom) and Afghanistan (Operation Enduring Freedom). The representative data are presented in bar chart format as shown in Figure 1-2. It can be inferred from the graph that IEDs alone contributed 36% of extremity injuries, while gunshots and grenades had accounted for 16% each.
Figure 1-1: Pie chart highlights the percentage of USA military loss due to landmine explosions in past warfare (Bird, 2001).
In addition to traumas to extremities, soft tissue injuries were reported among casualties. The soft tissue injuries related to IED explosions accounted for 53%, while associated fracture was 26%.
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Figure 1-2: Contribution of different causative agents towards extremities injuries caused during OIF & OEF (Owens et al., 2007).
Zouris et al.(2006) in their paper on combat casualties reported that landmines were the main causative agent for extremity injuries in Operation Iraqi Freedom (OIF). The authors considered 279 US Marines and soldier personnel deployed in the field for their study. Of the total reported casualties, landmines alone produced 79% of the injuries to the lower extremities whereas IEDs mainly contributed to upper extremity trauma, accounting for 36%. Table 1-1 presents a brief overview of the contribution of modern weapon’s to different orthopedic injuries. Therefore, it could be concluded that lower extremities are vulnerable to landmine explosion and upper extremities to IED explosive. The injuries sustained by the soldiers due to the landmines and IED explosions prevented them from performing an immediate action (Mckay, 2010).
Table 1-1: Injury location for wounded personnel due to different causative agents during OIF-I (Zouris et al., 2006).
Region (%) IED Landmine Mortar RPG Shrapnel Small Arms Total
Back 0 0 3.3 2.5 1.7 1.2 1.5 Lower extremities 28.2 78.8 33.3 25.9 29.3 31.7 34.4 Upper extremities 35.9 12.1 36.7 33.3 27.6 42.7 33.1 Pelvis 2.6 0 6.7 2.5 1.7 2.4 2.5 others 33.3 9.1 20 35.8 39.7 22 28.5 Total 100 100 100 100 100 100 100
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Other than injuries to the extremities, the spine is the second most vulnerable body region to experience high blast load under vertical loading. Schoenfeld et al. (2012) published a detailed review of combat-related spine injuries observed in 20th century Korean, Vietnam, and Gulf wars. The spine trauma accounted for only 1% of total combat causalities (Schoenfeld et al., 2012b). With the increase in the use of IEDs in recent wars, the probability of spine injuries has increased. The compressive load for the tibia was found to be much higher compared to the load experienced by the spine and the neck region (TR-HFM-090, 2007), shown in Figure 1-3. Next, the lumbar spine is reported to experience compressive vertical thrust of 4-5 kN.
Figure 1-3: Axial compressive force measured in the tibia, lumbar spine and upper neck for a vertical load caused by underbelly blast impact (TR-HFM-090, 2007).
In addition, to the vertical thrust through the seat pan, a considerable magnitude of mechanical load is transferred to the pelvis and the torso from the lower extremities due to floor intrusion. Figure 1-4 shows the +Gz vertical load transmission pathway through the body. Even though the lower extremities receive maximum compression force, the severity of the injury is high in the pelvis and the lower spine regions. The pelvis is home to many vital organs. The intervertebral and the sacroiliac joints are one of the most complex articulations of the body. Unlike the lower
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extremity bones, the pelvic innominate bones lacks soft tissue covering; hence, the lower spine receives a major portion of the seat load through the pelvic girdle. The complex morphology of the lumbar spine and the pelvis make this region of the body more vulnerable to UBB kinds of impact.
Figure 1-4: Pictorial representation of the load transmission paths subjected to vertical load under a UBB event (Ramasamy et al., 2009).
Comstock et al. (2011) in their investigation of Canadian warfare injury data states that the thoracolumbar spine region is more prone to field related spinal injuries. With reference to combat associated spinal trauma database, the authors reported that IEDs alone contributed about 57% of combat-related injuries followed by non-IEDs (23%) and blunt trauma related injuries (20%) in the live theater. Moreover, out of 372 injured soldiers from the Afghanistan war, 8% sustained at least one spine fracture. Further, 22 of 29 sustained a spinal injury due to an IED explosion. Among this spinal trauma, seven of them sustained stable fractures, nine unstable fractures, and six cases were unknown. Lumbar fracture was the most common battlefield spine injury.
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In another study, Ragel et al. (2009) reviewed the spinal injuries of the soldiers deployed in the Afghan war. They reported that among the spinal injuries sustained by mounted soldiers due to IED explosions, 38% of the fractures were noted in the thoracolumbar region alone. In addition, three soldiers sustained multiple vertebral injuries which included chance and burst fractures. These injuries were analyzed to occur when the spine, particularly the thoracolumbar region, undergoes hyperflexion compression. The authors pointed out that this behavior of the spine was found to be similar to the chance fracture mechanism sustained by fighter pilots during the ejection phase. The authors also mention that in an automotive accident the probability of chance and thoracolumbar fracture is less than 0.15% and 2.5%, respectively, while in UBB impact, the incidence of these fractures is 1.82% and 42% respectively.
Warfield-related pelvic trauma due to penetrating injury is fatal compared to blunt injuries. Bailey and Stinner et al. (2011) conducted an investigation on pelvic fracture and the related injuries sustained by soldiers, who either died from those wounds or were killed in action during operations in Iraq and Afghanistan. They selected 91 pelvis injury fatalities; of those 63 were mounted and 18 dismounted at the time of injury. Out of these total causalities, 66% of the injuries were grouped as penetrating injuries, whereas 34% were blunt injuries. Figure 1-5 shows that blast explosion is the major cause of pelvis fracture accounting for 74% of the total, whereas gunshot wounds and motor vehicle collisions were 15% and 4.5%, respectively (Bailey et al., 2011).
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Figure 1-5: Pie chart of the percentage of pelvic injury caused by different weapons (Bailey et al., 2011). Schoenfeld et al. (2012a) summarized the spinal injury sustained by soldiers deployed in Operation Iraq Freedom. They found that in the 15 month war, approximately 29 soldiers suffered from combat-related spinal trauma, accounting for 7.4% of the total combat casualties. Further, the blast mechanism produced 65% of the blunt trauma to the spine. 21% of those spine traumas were classified as closed fractures, whereas 7% were open fractures. Most of these combat-related spinal injuries were witnessed in the lumbar and cervical regions. The authors added that 7.4% of spinal trauma casualties in Operation Iraq Freedom (OIF) warfare were noted to be the highest in American warfare history, as shown in Figure 1-6. Other than load-bearing functionality, the lumbar region makes the upper torso flexible, bending forward, backward and sideward. Compare to the stiff thoracic spine, the lumbar spine is mobile, making the thoracolumbar junction more prone to fracture (Possley et al., 2012).
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Figure 1-6: Bar chart of the percentage of spine trauma in warfare injuries. *BCT- Brigade combat team deployed in OIF (Schoenfeld et al., 2012a).
1.3 Positioning of the occupant
The orientation of the occupants in an armored vehicle differs based on the duties assigned to them. Most infantry vehicles consist of a driver, a commander, gunner, and passengers. Except for the gunner, all other occupants’ feet rest directly on the floorboard and are in a seated posture. However, the gunner during a combat operation stands on an elevated platform; during a non-combat operation, he might sit on the seat. Based on the design of the vehicle, the seat arrangements of the occupants inside the vehicle differ in front or side facing. Also, the restraining and seat system differs from vehicle to vehicle. Figure 1-4 presents the flow of energy through the body during an underbody blast event. The orientation of the occupant during an impact plays a significant role in vertical load transfer through the body. However, for double site impacts such as those in UBB impact the high kinetic energy is transferred through both the feet and the buttocks. The pelvis tilt, the angle between the thigh and trunk, and the orientation
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of the sacrum with respect to the seat bottom are the primary parameters influencing the pattern of injury (Harrison et al., 1999). Schoberth and Hegemann (1962) examined the effect of the orientation of the pelvis on the body’s center on gravity shift. The representative orientation of the pelvis and the corresponding shift in the center of gravity are tabulated in Table 1-2 and shown in Figure 1-7. The authors stated that for a normal sitting posture, the angle at the hip, knee and ankle joints is (90-90-90). The ischial tuberosity is the point of support for a normal sitting posture accompanied by the posterior pelvis tilt. They further reported that most people prefer to sit in a relaxed state by tilting the pelvis more posteriorly. In this position, the lumbar spine tends to be straight or slight in convex with respect to the seat back. Moreover, the center of gravity of body mass while sitting is noted to shift dorsally from ischial tuberosity to the ischial lesser arch with an increase in the posterior tilt. Another important sitting position parameter is the thigh-trunk angle, which also plays a significant role in producing injury to the occupant exposed to UBB impact. Keegan and Omaha, (1953) studied the effect of the thigh-trunk angle during pelvic tilt and corresponding lumbar spine curvature, as shown in Figure 1-8. They pointed out that with the decrease in the thigh-trunk angle, the pelvis tilts posteriorly accompanied by the kyphosis of the mobile lumbar spine. The posture with the torso-femur angle at 90 degrees represents the sitting position of the occupant in a vehicle. During an underbody blast event, the lower spine and upper leg flex towards each other, reducing the torso-femur angle. Reduction in the angle is accompanied by posterior pelvic tilt, leading to more stress on the thoracolumbar spine (KEEGAN and Omaha, 1953).
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Table 1-2: The orientation of the pelvis, corresponding CG and body weight transferred with respect to sitting posture (Schoberth and Hegemann, 1962).
Sitting posture Orientation of pelvis Center of gravity location
Body weight transferred Anterior (A, B) Forward rotation of the pelvis. Flexing spine without much rotation of pelvis
In front of the ischial tuberosity
Feet transmit more than 25%
Middle (C) Neutral or normal position
Above ischial tuberosity
Feet transmit 25% Posterior (D) Extension rotation
of the pelvis.
Kyphosis of the spine
Above/behind ischial tuberosity
Feet transmit less than 25%
Figure 1-7: Shift in the body’s center of gravity based on occupants sitting posture (Schoberth and Hegemann, 1962).
During an underbody blast, the occupant experiences two separate +Gz accelerations (Figure 1-4) (Ramasamy et al., 2009). The first input is at the feet of the occupant due to the intrusion of the floor-plate. The second input is at the occupant’s buttocks due to the vehicle’s upward acceleration. The feet acceleration precedes the seat acceleration by a few milliseconds. Sacral and pelvic injuries are mostly due to seat acceleration. This combination of short duration high- rate loading can results in complex injury patterns.
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CHAPTER 2
THEATER INJURY CASE STUDY AND BLAST BIOMECHANICS
The increased use of landmines/ IEDs by insurgents in Iraq and Afghanistan confirms that these are the signature weapon of the future battlefield (Owens et al., 2007; Ramasamy et al., 2009). Blast events due to landmine explosion have caused major losses to the US military in terms of both human and wealth resource. Injuries related to IED explosions are entirely different and mores server compared to other modern weaponry (Bailey et al., 2011; Owens et al., 2007; Zouris et al., 2006). Although the duration of mines/IED explosion events occurs in a few milliseconds, the kinetic energy generated by the detonation is significant enough to disintegrate the vehicle (Wang et al., 2001). Simultaneously, occupants of the vehicle experience severe orthopedic and soft tissue injuries. The injury mechanisms due to blast impact could well be understood by reviewing IED associated injury case studies of soldiers subjected to underbody blast impact and examining the basic concept of the physics behind the blast event.
1.4 Case study from literature
There are only a few live theater IED-associated spinal and pelvic trauma case studies in the literature. Soldiers returning with tertiary blast injuries from the recent battlefield are very complex to operate it back to normal state. No such injury pattern has ever been reported among the civilian population.
One such multiple and complex military trauma due to an IED was seen at Walter Reed Military Hospital. Kang et al. (2012a) illustrated a classic example of a comminuted Zone III sacral fracture along with a bilateral sacroiliac joint rupture and three transverse process fractures sustained by an on-duty soldier exposed to an IED explosion. Moreover, the solider also received
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associated injuries that included a bilateral trans-tibial amputation, a left acetabular fracture, multiple rib fractures and organ injuries.
In another case study witnessed at the same hospital reported that an active soldier exposed to similar environmental conditions sustained a stable Zone II fracture at S1/S2 (Cody et al., 2012). The associated injuries included a left L5 transverse process fracture and a facet fracture with L5/S1 retrolisthesis.
A similar IED induced trauma active solider was received at the University of Health Science, Maryland. He sustained L5 and L4 burst and compression fractures, respectively, along with a posterior ligament ruptures (Kang et al., 2012b). The associated injuries included trans-femoral amputation and facial fractures.
To understand the biomechanical response and the mechanism invovled in these live theater injuries, cadaver testing under a controlled environment has been performed. Bailey and Christopher et al. performed PMHS testing under a laboratory setup using a sled system (Bailey et al., 2013). They examined the pelvis and lower extremity mechanical response when exposed to UBB loading conditions. Five whole-body cadavers were subjected to impactor velocity ranging from 7.5m/s to 14m/s over a 3 millisecond interval. The corresponding pelvis and tibia threshold was measured to be 300g and 600 g respectively. The surrogates sustained a combination of pelvis, spine, and lower leg injuries. The authors also determined the Hybrid III dummy response under similar loading conditions. They found that both the loading rate and pelvis jerk to be higher in the dummy compared to cadaver testing. Further they reported the inexactitudes of the automotive dummies in replicating cadaver response under UBB conditions.
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1.5 Effect of blast event on a vehicle and its occupant
Triggered IEDs and landmines irrespective of their state (liquid, solid or gaseous) undergo a quick chemical reaction, producing a pressurized gaseous product. This compressed gas expands rapidly in the surrounding area, and its volume increases to about 105 times the atmosphere volume. Therefore, gas molecules are accelerated, leading to the formation of a shock wave which then propagates at a velocity of 1,000 m/s (Stuhmiller et al., 1991). For Trinitrotoluene composed IEDs, detonation velocity was determined to be 7,000 m/s (Schardin, 1950). The resulting wave pattern disturbs the state of the surrounding gas molecules. In turn, temperature rises in the range of 2000 to 6000 C. The density and pressure of the blast wave produces a distorted region (Ramasamy et al., 2011). The static or shock wave followed by a rapid motion of gas molecules produces blast wind. Human or animal tissue exposed to or lying in this environment will yield serious blast injuries.
The complexity of the detonation waveform and the energy liberated from the explosion determines the intensity of the wound caused. Underbody blast events and IED/ landmine explosion procedures follow three major phases (Ramasamy et al., 2011). First, the blast wave produced due to the detonation interacts with the soil. Second, the highly compressed shock wave fractures the surface of the soil, liberating the gas molecules which further hit the base of the vehicle. The incident wave is reflected from the vehicle base toward the center of the explosive. Hence, the reflected wave multiplies with the incident wave, generating a highly compressed region between the soil surface and the base of the vehicle.
Third, these pressurized gas molecules and soil eject due to the gas expansion during detonation interact with the vehicle floor instigating local deformation and fracture of the floorboard. This
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ruptured base of the vehicle allows the pressurized shock wave to penetrate the occupant compartment resulting in injuries to the lower extremities in the first phase. In the second phase, the whole vehicle is accelerated vertically, producing injuries to the upper leg, pelvis, and spine.
Based on the landmine/IED explosion linked causality database from war in Iraq and Afghanistan, blast-related injuries are classified into four main categories: primary, secondary, tertiary and quaternary (Ramasamy et al., 2009). These will be explained below.
The initial shockwave after detonation rapidly increases the surrounding pressure. The injury produced under such an atmosphere is termed as a primary blast injury (Ramasamy et al., 2009). The blast wave does not accelerate the body. However, the sudden change in pressure causes serious injuries to hollow organs such as the GI tract and lungs. The primary blast wave is also observed to produce mild traumatic brain injury (mTBI) (Taber et al., 2006; Warden, 2006; Warden et al., 2009). The exact mechanism involved in mild TBI is yet to be investigated. The severity of the injury depends on the distance of the body from the explosive site and also the amount of explosive used for detonation (Kang et al., 2012c).
Secondary injuries occur due to the accelerated explosive or nearby compartment fragments. The pressurized detonation product transfers the energy and momentum to the vehicle body that in turns accelerates the debris. Based on the kinetic energy of the wreckage and its material property, injury severity varies. Lower extremity fractures are the most common trauma observed in secondary blast injuries.
Both the local and global effects resulting from the explosion yield tertiary injury. The highly pressurized shock wave and accelerated soil debris accelerate the vehicle. The accelerated
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occupant compartment causes the occupant to collide with the interior surface, and occupants thrown from the seat, resulting in significant injuries to the upper legs, pelvis, and spine (Ramasamy et al., 2009). Most of these injuries are related to the long bones and vertebral body fractures. Head and facial injuries are other common wounds seen in such an impact (DePalma et al., 2005; Xydakis et al., 2005). The magnitude of the explosive and the mass of the vehicle determines the vertical acceleration of the vehicle in the air (Kang et al., 2012c). After reaching a certain height, the vehicle drops down due to the action of gravity. The load due to gravity also acts upon the passenger in the vehicle. However, this load is insignificant compared to the vertical thrust (Mckay, 2010). Local deformation of the floorboard produces serious injuries and fracture to the lower extremities. Most UBB-associated trauma could be classified as tertiary blast injuries. Injuries due to thermal burns and the aftermath of detonation together constitute quaternary injuries.
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CHAPTER 3 ANATOMY AND BONE MINERAL DENSITY
During the tertiary phase of an underbody blast event, the blast overpressure and the blast wind accelerate the vehicle and its passengers above the ground. Due to the intrusion of the floor plate, the lower extremity region such as the ankle bones, tibia, femur, and patella are loaded with the high vertical load. Followed by floor deformation, the seat is translated vertically upward resulting in pelvis and lower spine injuries. The increase in area/volume and the load carrying capacity of the lumbar vertebrae saves the superior spinal segments from injury under higher vertical acceleration (Yoganandan et al., 2013). To understand the orthopedic injury mechanism of different regions, the anatomy of the chief bones must be analyzed.
1.6 Spine
The vertebral column of a human, which lies medially to the posterior part of the trunk, is one of the most complex musculoskeletal structures. It runs all the way from the base of the skull to the pelvic girdle. Its primary function is to protect the spinal cord and support the head, neck, upper extremities and trunk. It also transfers the load from the trunk to the pelvis. There are in total 33 vertebrate: 7 cervical, 12 thoracic, 5 lumbar, 5 sacral, 4 coccyx. The first 24 are well defined and articulate with successive vertebrae, and the remaining 9 are fused to form the posterior frame of the pelvic girdle (Gray et al., 1973). This bony spinal column forms two curvatures, namely kyphotic (thoracic and sacral curve) and lordosis (cervical and lumbar), as shown in Figure 3-2. These curves help with balancing the body weight and walking. In addition, the 22 fibrocartilage discs between adjacent vertebrae form synovial joints, which allow the movement of the spine in all three anatomical planes and act as a shock absorber. No such disk is found between the skull and C1 or between C1 and C2. The vertebral disks along with the abdomen and back muscles
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stabilize the spinal column, while the anterior and posterior muscles of the spine provide the force required for the flexion and extension movements of the trunk, respectively (Moore et al., 2006). These muscles work in unison as well and provide the rotational ability to the spine. The vertebral column is the main channel for transferring the load from caudal to cranial and vice-versa (Gray et al., 1973). Therefore, the load experienced by the lower extremities, pelvic or head ultimately leads to an indirect impact on the vertebral column.
Thoracic and lumbar vertebral fractures are common, under a UBB-induced spinal trauma (Comstock et al., 2011; Possley et al., 2012; Ragel et al., 2009). The thoracic vertebrae make up the central part and occupy a larger portion of the spinal column (Netter, 2010). These vertebrae have long and almost horizontal spinous processes. Furthermore, the facet is more vertical and oriented in the coronal plane. Also, the body diameter increases from T1 to T12, with the T1 centrum resembling the cervical vertebrae body and T12 as L1, respectively. In high +Gz acceleration, T12 is prone to stress/ compression related injury. Of all the 33 vertebrae, the lumbar are the largest bones of the spine (Gray et al., 1973). In addition, these vertebraes have a flat superior, and inferior end plates, which make them to bear the load applied along the axis. Moreover, with the curved and vertical facets, the lumbar vertebrae have the ability to withstand the shear load. Unlike, the thoracic vertebrae, the lumbar do not have a prominent spinous process. The transverse and spinous processes of the lumbar vertebrae function like a lever, enhancing the function of the muscles attached to them (Bogduk, 2005). Figure 3-1 shows a pictorial image of lumbar and thoracic vertebral body.
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Figure 0-1: A pictorial comparison between the lumbar and thoracic vertebral bodies (Gray et al., 1973).
Figure 0-2: Anatomy of the Spine (Gray et al., 1973).
1.7 Pelvis
Each of the two innominate hemi-pelvis bones consist of three sub-bones: the ilium, ischium and pubis. Each fuses at the acetabulum, forming the anterior and lateral walls of the pelvic girdle whereas, the fused sacral and coccyx vertebrae form the posterior part of the girdle (Gray et al., 1973). The sacrum is a wedge between the hip bones which is attached to the ilium bone of the
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hip by interosseous ligament, forming the sacroiliac (SI) joint (Netter, 2010). This connection of the sacrum to the innominate bones along with the symphysis of the pubic bone provides a ring appearance to the pelvic girdle. The pelvic ring transmits the load from the vertebral column towards the lower extremities and vice versa (Moore et al., 2006). In addition, the upper body weight converges to the femoral neck through the sacroiliac joint. A high-energy impact such as UBB event disrupts the pelvic bone as well as abdominal organs. An anterior anatomical view of pelvic girdle is presented graphically in Figure 3-3.
Figure 0-3: Anatomy of the Pelvis (Gray et al., 1973).
1.7.1 Sacrum
In a UBB scenario, mounted occupants experience two separate vertical accelerations. One is at the feet due to floor plate intrusion, and the other is at the pelvis due to seat pan acceleration. The sacrococcygeal segment of the spine comes in direct contact with seat pan. High +Gz seat acceleration contributes to severe pelvic and sacral fracture. The Sacrum articulates with the fifth lumbar vertebrae cranially whereas its apex articulates with the coccyx (Gray et al., 1973). In addition, it forms the sacroiliac joint with the hip bones, which are held together by a sacral-iliac
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ligament. Five fused curved sacral vertebrae project forward, resulting in the sacrovertebral angle with the last lumbar vertebrae (Netter, 2010). As the central region is curved and directed backward, this increases the capacity of the pelvic cavity. The size of the sacral vertebrae decreases from top to bottom. Each ridge shown in Figure 3-5 ends in a sacral foramina, through which the sacral nerve pass (Gray et al., 1973). These lateral foramina are a source of stress concentration, and a vertical fracture could result through these foramina.
Figure 0-4: Sacrum frontal view (Gray et al., 1973).
Furthmore, the posterior surface of the sacrum is narrower and more convex then the anterior portion. It consists of incomplete spinous processes and laminae. Other than the SI ligament, sacrotuberous and sacrospinous ligaments play an essential role in locomotion and maintaining the stability of the pelvis (Moore et al., 2006). The sacrospinous ligament resists external rotation of the ischium, while the sacrotuberous ligament resists vertical loads by pushing the pelvis down with respect to the sacrum. This action of the ligament results in serious injuries to both the pelvis and sacrum during vertical seat pan acceleration, such as in an IED detonation. In UBB events, both vertical and transverse sacral fractures have been observed to be common among mounted soldier (Cody et al., 2012; Kang et al., 2012a).
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1.8 Ligaments of spine and pelvis
The pelvis lacks soft tissue. Hence, it is deprived of inherent stability. The ligament and muscles of the spine and pelvis are essential parameters for pelvic stability. The pubic rami and symphysis joint together act as a strut and prevents the pelvis from collapsing anteriorly, while the SI complex provides the stability to the posterior structure (Gray et al., 1973). Analyzing the pelvic injury mechanism requires careful assessment of both the magnitude and direction of the load as well as the orientation of the pelvis at the time of impact. To some extent individual upper body mass plays an important role in defining the severity of the injury. The femoral neck receives the torso weight through the SI complex (Gray et al., 1973). Therefore, a larger body mass would destabilize the hip joint resulting in lower back and limb pain. Table 3-2 presents the ligaments involved in stabilizing pelvic structural integrity. Most of the connecting and pelvic floor ligaments work as one group to maintain pelvic ring structure stability. The posterior ligaments of the spine and the pelvis together form a tension band that resists the deforming forces (Tile et al., 2003), while the SI, iliolumbar and the sacrospinous ligament resist the transverse rotational force and the vertical ligaments running along the pelvis resist the shearing force in the longitudinal direction.
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Table 0-1: A brief summary of pelvis-floor ligament anatomical location and physiological function. (These ligaments have been hypothesized to play a critical role in pelvis fracture mechanism under, a UBB type environment) (Gray et al., 1973; Tile et al., 2003).
Ligament Anatomical location Function
Iliolumbar (IL) Transverse process of L5 to iliac crest
Strength the lumbar-sacral joint. Interosseous Sacroiliac lateral surface of the ilium to the
lateral surface of the sacrum
Resist abduction of SI joint. Resist anterior displacement of the sacrum.
Anterior Sacroiliacs (ASI) Anterior aspect of sacrum to ilium Resist external rotation and shearing force
Posterior Sacroiliacs (PSI)
Posterior superior iliac spine (PSIS) to lateral surface of sacrum
tension band of pelvic ring and stabilize the posterior pelvis Sacrospinous Connects the lateral edge of the
sacrum to ischial spine
Resist external rotation force Sacrotuberous Connects the sacrum to ischial
tuberosity
Resist shearing rotary force. Pubic Symphysis Joins the lateral aspect pubic bone Along the pubic rami, it acts as a
strut and maintains the pelvis stability anteriorly.
1.9 Femur
The femur is the longest and strongest cylindrical bone in the body, shown in Figure 3-4. It joins as well as transfers the load from the pelvic girdle to the lower leg. The globular head articulates with the acetabulum to form a ball and socket joint (Huelke, 1986). In addition, a flat pyramidal portion of the bone termed as the femoral neck forms a 125 degree angle with the shaft. Furthermore, the shaft runs down to form medial and lateral condyles that articulate with the tibia condyle to form the knee joint. A large group of muscles called the quadriceps femora anchors to the lateral and anterior region of the femur (Netter, 2010). These muscles act as an extensor for the knee joint and stabilize the patella bone in position. This muscle is essential for walking, running and squatting.
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Figure 0-5: Anterior and posterior view of the femur (Gray et al., 1973). 1.10 Bone mineral density(BMD)
The severity of a fracture depends on both the amount of load experienced by the skeleton structure due to impact and the bone strength of the occupants involved (Melton et al., 1997). Although soldiers have good physique and undergo similar training, the bone mineral content of each differs. The BMD differs based on gender, age, race, and ethnicity. From beginning of the civil war the contribution of black Americans in the war field has been significant, as shown in Figure 3-6 (Segal and Segal, 2004). Later due to a large number of immigrants, people of different races and ethnicities came forward to join the US military. Table 3-2 shows active duty service officers based on their rank, race, and ethnicity. The authors state that all four US military branches consist of people from different backgrounds (Segal and Segal, 2004). Futher, the posting of these personnel is determined purely by their performance and skill and not on their race. The US military consists of one of the most diverse populations in the world (Segal and Segal, 2004).
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Bone is made up of minerals, principally calcium hydroxyapatite, embedded in type 1 collagen and a specialized protein that makes up the bone matrix (Cummings et al., 2002). Mineral content is measured using bone densitometry. Calcium is capable of absorbing more radiation compared to other minerals. The higher the mineral content, the darker the image will be, indicating a larger amount of calcium content at the particular point of the bone.
Figure 0-6: Active duty military and civilians by race and ethnicity, 2002 (Segal and Segal, 2004). Table 0-2: Active –service duty officers by rank, service and race/ethnicity (Segal and Segal, 2004).
Rank Army Navy Marines Air Force
Whi te B lac k H ispa n ic Whi te B lac k H ispa n ic Whi te B lac k H ispa n ic Whi te B lac k H ispa n ic All the officers (%100) 100 100 100 100 100 100 100 100 100 100 100 100 Company grade 57 61 71 58 69 73 62 71 79 57 63 64 Field grade 42 39 29 42 30 27 37 29 20 43 37 36 General 1 -- -- -- -- -- -- -- -- -- -- --
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Bone mineral density (BMD) is defined as the ratio of bone mineral content to the area of the bone that is being analyzed (Cummings et al., 2002). Clinical bone densitometry results are defined in terms of T-scores and Z-scores. The standard deviation (SD) obtained by comparing patients’ densitometry with the BMD of young, healthy adults is termed as the T-score (NIH, 2012). In contrast, if the patient mineral content is compared to the BMD of the same age group, the obtained SD value is termed as the z-score. An SD below -2.0 indicates the BMD of the occupant is low compared to the general population. A Z-score measurement sometimes could be misleading since the SD value is predicted based on comparison with the same age group (NIH, 2012). Table 3-8 shows the World Health Organization’s proposed T-scores and Z-scores.
Table 0-3: World Health Organization’s definitions based on Bone Density Levels (T-score) (NIH, 2012). According to World Health Organization
Level Definition
Normal Bone density is within 1 SD (+1 or −1) of the young adult mean.
Low bone mass (osteopenia) Bone density is between 1 and 2.5 SD below the young adult mean (−1 to −2.5 SD).
Osteoporosis Bone density is 2.5 SD or more below the young adult mean (−2.5 SD or lower).
Severe (established) osteoporosis Bone density is more than 2.5 SD below the young adult mean, and there have been one or more osteoporotic fractures.
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Table 0-4: Osteoporosis attribution probability by fracture type, race/ethnicity and age (Melton et al., 1997).
45-64 years 65-84 years ≥ 84 years
Site Median attribution, probability (range) Validity rank Median attribution, probability (range) Validity rank Median attribution, probability (range) Validity rank Whi te po pul at ion Hip 0.60 (0.10-0.70) 2.2 0.80(0.60-0.95) 1.8 0.85(0.80-0.95) 1.7 Spine 0.70 (0.50-0.90) 2.2 0.90(0.50-0.95) 1.8 0.90(0.60-0.95) 1.8 Forearm 0.40 (0.05-0.50) 2.5 0.45(0.15-0.60) 2.3 0.45(0.30-0.60) 2.2 Other sites 0.15 (0.05-0.30) 2.7 0.30(0.20-0.40) 2.7 0.45(0.30-0.50) 2.7 B lac k p opul at ion Hip 0.30(0.05-0.65) 2.8 0.65(0.10-0.85) 2.3 0.75(0.25-0.90) 2.3 Spine 0.55(0.30-0.40) 3 0.75(0.30-0.90) 2.5 0.85(0.30-0.95) 2.3 Forearm 0.20(0.05-0.40) 2.7 0.30(0.10-0.50) 2.8 0.35(0.20-0.50) 2.8 Other sites 0.15(0.05-0.20) 3.5 0.15(0.05-0.30) 3.5 0.25(0.15-0.40) 3.5 O ther R ac e popul at ion Hip 0.55(0.10-0.65) 3.2 0.75(0.15-0.90) 3 0.85(0.30-0.95) 3 Spine 0.60(0.30-0.80) 3.2 0.75(0.40-0.90) 3 0.85(0.50-0.95) 3 Forearm 0.30(0.30-0.55) 3 0.35(0.15-0.50) 3 0.40(0.30-0.50) 3 Other sites 0.15(0.10-0.30) 3.3 0.20(0.10-0.40) 3.3 0.30(0.20-0.50) 3.3
The age, gender, race and ethnicity of the patient are essential for defining the bone mineral density of the occupant. Based on the standard deviation value the World Health Organization has defined a certain range for predicting the occurrence of osteoporosis and osteopenia, as presented in Table 3-4. It represents the variation in the standard deviation of different races and ethnicity groups that determines the probability for osteoporosis leading to bone fracture (Melton et al., 1997). Furthermore, the fracture rate is higher in Caucasian men than black men, whereas the other groups lie between these ranges. In addition, age is also an important parameter for determining the mineral content of bone. For example, consider a data from Table 3-4, the probability of hip bone fracture for 45 years old white man lies in the range of 0.10 to 0.70, while
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it increases to 0.60 to 0.95 for 65-year-old man of the same race. Moreover, when the same parameter was considered for black men of age 45, the chance of getting a hip fracture was noted to be low (0.05-0.65). Authors mentions that white and asian groups are prone to fracture since their bone mass is comparatively less dense than for black men (Melton et al., 1997). Therefore, based on race and ethnicity the reference value for determining a patient's T-score and Z-score will differ. Gender consideration is also an important parameter in determining bone mineral content. However, for an UBB event study, male gender is considered since most of the mounted soldiers in an army tank are male. Hence, the bone mineral density value for female the gender is not taken into consideration for this study. In an UBB impact, high loads are experienced by the bone leading to a fracture. Unlike the magnitude of the load and duration of the event, bone mineral density or bone strength plays a minimal role in understanding the injury and biomechanical response of the occupant. However, mineral density measurement is one of the important criteria in the specimen selection procedure. In the current study, it is used as a tool to reject a specimen with osteoporosis and osteopenia conditions prior to acquiring a body from a vendor. All the cadaveric BMD were measured using the Dual X-ray Absorptiometry technique.
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CHAPTER 4
CADAVER TESTING AND DATA PROCESSING METHODS
UBB-associated casualties have recently increased due to the wars in Iraq and Afghanistan. Such complex injury patterns are rarely seen among civilians in car crashes. Researchers and the military are performing collaborative studies with armored vehicle manufacturers to mitigate these injuries (Bailey et al., 2015; Bosch et al., 2014; Kargus et al., 2008). However, automotive dummies have not yet been developed for evaluating occupant response subjected to vertical loading conditions. In addition, the currently available biomechanical response data are not suitable for designing a new biofidelic dummy. Therefore, to obtain the characteristic whole body response from UBB impact conditions, the military is funding cadaver testing under a controlled environment. Whole body post mortem human subject (PMHS) testing under laboratory conditions will provide a better understanding of the mechanical response of the body subjected to vertical load and will possibly explain combat-linked injuries due to IED/landmine explosions. Finally, the biomechanical response colliders developed from cadaver testing could be implemented in dummy design for future studies and live fire testing.
The current study thesis on investigating the biomechanical response and corresponding injuries sustained by the test subject for given UBB loading condition and also on the development of the methodology for this testing. Using the modified Wayne Horizontal Acceleration Mechanism (WHAM) III sled system, two whole body specimens were subjected to simulated UBB impact loading conditions under a controlled laboratory set up. A detailed overview of the sled system is discussed in the following section.
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1.11 Horizontal sled system
The WHAM III sled deck measures 4 x 2 m and is capable of producing deceleration pulse controlled by a hydraulic decelerating mechanism. A set of two parallel rails were mounted to the sled deck. A vertical rigid seat was reclined (rotated along the global Y axis) so that the seat back was parallel to sled deck and rails. Linear roller bearings were attached to the rear aspect of the seat fixture and it was positioned onto the rails and coupled to the rails with retention brackets. The roller bearings and the retention bracket assemblies illustrated in Figure 4-1 allowed the rigid seat to move freely on the rails during the impact event.
Figure 0-1: A lateral view of the horizontal sled system with the occupant positioned on the seat with a 5-point belt.
A movable foot floor assembly capable of producing independent foot floor pulse consists of a 0.406 m by 0.406 m by 0.0063 m floor plate, a linear bearing system, a 0.05 m diameter cylindrical shaft, and an elastomer. The cylindrical shaft was rigidly connected to the floor plate on the occupant end (Figure 4-1). The linear bearing system was mounted to the seat fixture and it was designed to constrain the motion of the shaft to move along the anatomical Z direction
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only. On the barrier end, the cylindrical shaft was connected to an elastomer to allow impact to the rigid barrier independent of the seat motion (Figures 4-1 and 4-2). Before the test, the movable floor system was adjusted along the anatomical X axis so that the line joining the centers of the occupant’s feet was orthogonal to and at the level of the cylindrical shaft. The sled was accelerated to a target velocity before the external force was released for the last 9.75 m of travel. Two large capacity snubber pistons (model RCOS 3X 12 BS 04 Efdyn Inc., OK) which were mounted to the barrier (Figure 4-2) were used to decelerate the WHAM III sled while allowing the seat system to continue the forward motion. Four pre-crushed aluminum honeycomb blocks which were attached to the barrier (Figures 4-1 and 4-2) were used to arrest the seat motion, while allowing the foot floor- shaft-elastomer assembly to impact the barrier directly. The total cross-sectional area and crush strength of the pre-crushed aluminum honeycomb blocks determined the deceleration pulse length. By controlling these two parameters, the crash pulse of the seat can be adjusted.
At the foot floor plate, the elastomer attached to the cylindrical shaft first deforms upon impact and then returns the stored energy and pushes the floor plate towards the occupant's feet. The stiffness of the elastomer attached to the cylindrical shaft determined the floor pulse magnitude, while the time to peak velocity for the floor pulse was adjusted by adding the weight plates to the barrier end of the shaft. Based on the weight added, the energy absorbed by the plates delays the time to peak for the floor pulse. During the test preparation phase effort were made to avoid any gap between the elastomer and the barrier, and between the seat and the pre-crushed honey comb blocks. This was carried out in order to achieve same time of arrival for both the seat and floor pulse.
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Figure 0-2: Two large capacity snubbers mounted to the barrier to slow down the WHAM III. Four pre-crushed aluminum honeycomb blocks attached to the rigid barrier to produce short duration seat acceleration.
1.12 Data acquisition and camera setup
A slice Pro (Diversified Technical Systems, Inc., CA) data acquisition system was used for each test to record 98 channels of data, including eleven 6DX blocks, strain gauges, contact switches, seatbelt load cells, fixture velocity, and floor and seat accelerations. The mounting locations of these 98 channels are listed in Appendix B. All the channels were sampled at the rate of 500,000 samples per second with a 100 kHz anti-aliasing, multiple low pass Butterworth filter. The event was recorded using two NAC GX-1 cameras (NAC Image Technology, CA) each with a 24 mm Nikon lens. One camera recorded the lateral view of the event, while the other camera recorded an overhead (frontal) view. The lateral view camera was positioned 1.57 m from the right side edge of the seat. The overhead camera was mounted 1.77 m from the top of the floor plate. Each